U.S. patent application number 12/669556 was filed with the patent office on 2010-07-22 for x-ray generating apparatus and x-ray ct apparatus using the same.
This patent application is currently assigned to HITACHI MEDICAL CORPORATION. Invention is credited to Keiji Koyanagi, Yoshiaki Tsumuraya, Hironori Ueki.
Application Number | 20100183117 12/669556 |
Document ID | / |
Family ID | 40259748 |
Filed Date | 2010-07-22 |
United States Patent
Application |
20100183117 |
Kind Code |
A1 |
Tsumuraya; Yoshiaki ; et
al. |
July 22, 2010 |
X-RAY GENERATING APPARATUS AND X-RAY CT APPARATUS USING THE
SAME
Abstract
There are provided an X-ray generating apparatus capable of
switching X-ray beams of high energy and low energy to each other
at high speed, and an X-ray CT apparatus capable of performing
high-speed and high-quality multi-energy imaging by using the same.
The X-ray generating apparatus is constructed by an X-ray tube 9
having two anodes 200a, 200b, a rotational anode 204 for radiating
X-ray from an X-ray focal point by electron beams emitted from
filaments of these cathodes, and two grid electrodes 202a and 202b
for controlling emission of the electron beams, a tube voltage
generator 9a and a tube voltage controller 9d1 for controlling an
X-ray condition, a filament heater 9b and a tube current controller
9d2, a grid voltage generator 9c and a grid opening/closing
controller 9d3, and a grid switching unit 9e. High energy X-ray and
low energy X-ray are switched and emitted to an examinee every
adjacent projection angles, thereby collecting projection data.
Inventors: |
Tsumuraya; Yoshiaki; (Tokyo,
JP) ; Ueki; Hironori; (Tokyo, JP) ; Koyanagi;
Keiji; (Tokyo, JP) |
Correspondence
Address: |
COOPER & DUNHAM, LLP
30 Rockefeller Plaza, 20th Floor
NEW YORK
NY
10112
US
|
Assignee: |
HITACHI MEDICAL CORPORATION
TOKYO
JP
|
Family ID: |
40259748 |
Appl. No.: |
12/669556 |
Filed: |
July 18, 2008 |
PCT Filed: |
July 18, 2008 |
PCT NO: |
PCT/JP2008/063015 |
371 Date: |
January 18, 2010 |
Current U.S.
Class: |
378/9 ;
378/111 |
Current CPC
Class: |
H05G 1/58 20130101; H01J
2235/068 20130101; A61B 6/405 20130101; A61B 6/482 20130101; H01J
35/26 20130101; A61B 6/4021 20130101 |
Class at
Publication: |
378/9 ;
378/111 |
International
Class: |
A61B 6/03 20060101
A61B006/03; H05G 1/32 20060101 H05G001/32 |
Claims
1. An X-ray generating apparatus including an X-ray tube for
emitting X-ray, tube current control means configured to control
tube current of the X-ray tube, and X-ray control means configured
to control high-energy X-ray and low-energy X-ray by the tube
voltage control means, characterized in that the X-ray tube has
plural cathodes each of which has a filament, an anode opposed to
the plural cathodes, and grid electrodes each of which is
individually provided to every cathode to control discharge of an
electron beam emitted from the cathode, and comprises grid voltage
generating means configured to generate a voltage to be applied to
each grid electrode, and electron beam emission control means
configured to alternately apply the grid voltage generated in the
grid voltage generating means to each of the grid electrodes to
control the emission of the electron beam.
2. The X-ray generating apparatus according to claim 1, wherein the
X-ray tube is an X-ray tube for emitting plural electron beams from
the plural filaments and forming plural X-ray focal points on the
anode so that the X-ray focal points are spaced from one another at
a predetermined distance on the anode.
3. The X-ray generating apparatus according to claim 1, wherein the
X-ray tube further comprises electron beam deflecting means for
deflecting the directions of the electron beams generated from the
plural filaments.
4. The X-ray generating apparatus according to claim 3, wherein the
electron beam deflecting means has a deflection coil provided
between the anode and the plural cathodes and deflection current
supply means for supplying the deflection coil with current for
deflecting the directions of the electron beams.
5. The X-ray generating apparatus according to claim 1, wherein the
tube voltage control means has first tube voltage control means
configured to control a high tube voltage corresponding to the high
energy X-ray, and second tube voltage control means configured to
control a low tube voltage corresponding to the low energy X-ray,
and the tube current control means has first tube current control
means configured to control tube current corresponding to the high
energy X-ray, and second tube current control means configured to
control tube current corresponding to the low energy X-ray.
6. The X-ray generating apparatus according to claim 5, wherein the
tube current controlled by the second tube current control means is
larger than the tube current controlled by the first tube current
control means.
7. An X-ray CT apparatus that has a multi-energy imaging function,
and comprises an X-ray tube configured to irradiate X-ray to an
examinee, an X-ray detector configured to detect X-ray transmitted
through the examinee, scanner rotating means rotating around the
examinee while the X-ray tube and the X-ray detector are mounted
therein, X-ray control means configured to irradiate X-rays having
plural different energies emitted from the X-ray tube to the same
slice position of the examinee while switching the X-rays every
projection angle, and image reconstructing means configured to
reconstruct projection data detected by the X-ray detector to
obtain a CT image, characterized in that the X-ray tube comprises
plural cathodes each of which has a filament, an anode opposed to
the plural cathodes, and grid electrodes each of which is
individually provided every cathode to control emission of an
electron beam emitted from the cathode, wherein the X-ray control
means comprises tube current control means configured to heat the
cathode filaments of the X-ray tube and control the tube current
flowing between the anode and the cathodes, tube voltage control
means configured to control a tube voltage to be applied between
the anode and the cathodes, grid voltage generating means
configured to generate a voltage to be applied to each grid
electrode, and electron beam emission control means configured to
apply the grid voltages generated in the grid voltage generating
means to the respective grid electrodes while alternately switching
the grid voltages at every projection angle, thereby controlling
emission of the electron beams.
8. The X-ray CT apparatus according to claim 7, wherein the X-ray
tube is an X-ray tube for emitting plural electron beams from the
plural filaments and forming plural X-ray focal points on the anode
so that the X-ray focal points are spaced from one another at a
predetermined distance on the anode.
9. The X-ray CT apparatus according to claim 7, wherein the X-ray
tube further comprises electron beam deflecting means for
deflecting the directions of the electron beams generated from the
plural filaments.
10. The X-ray CT apparatus according to claim 9, wherein the
electron beam deflecting means has a deflection coil provided
between the anode and the plural cathodes and deflection current
supply means for supplying the deflection coil with current for
deflecting the directions of the electron beams.
11. The X-ray CT apparatus according to claim 7, wherein the tube
voltage control means has first tube voltage control means
configured to control a high tube voltage corresponding to the high
energy X-ray, and second tube voltage control means configured to
control a low tube voltage corresponding to the low energy X-ray,
and the tube current control means has first tube current control
means configured to control tube current corresponding to the high
energy X-ray, and second tube current control means configured to
control tube current corresponding to the low energy X-ray.
12. The X-ray CT apparatus according to claim 11, wherein the tube
current controlled by the second tube current control means is
larger than the tube current controlled by the first tube current
control means.
13. The X-ray CT apparatus according to claim 7, wherein the number
of projection angles for detecting the projection data is larger
than that under normal CT imaging.
Description
TECHNICAL FIELD
[0001] The present invention relates to an X-ray generating
apparatus, and particularly to an X-ray generating apparatus
suitable for multi-energy imaging, and an X-ray CT apparatus having
a function of performing multi-energy imaging by using the X-ray
generating apparatus.
BACKGROUND ART
[0002] As X-ray CT apparatuses are known a single slice type X-ray
CT apparatus for obtaining one tomogram by one X-ray exposure and a
multi-slice type X-ray CT apparatus for simultaneously obtaining
plural tomograms by one X-ray exposure.
[0003] According to the single slice type X-ray apparatus, an X-ray
detector having many X-ray detection elements arranged on a line,
that is, in a one-dimensional direction (channel direction) is
used, a fan beam, that is, a sectorial X-ray beam is emitted from
an X-ray tube to an examinee and X-ray which has passed through the
examinee is measured to obtain projection data of the examinee.
[0004] On the other hand, according to the multi-slice type X-ray
CT apparatus, a cone beam, that is, a conical or pyramidal X-ray
beam is emitted from an X-ray tube to an examinee, and X-ray which
has passed through the examinee is measured by an X-ray detector
having many X-ray detection elements arranged in a two-dimensional
direction (channel direction and line direction), thereby obtaining
projection data.
[0005] In both the X-ray CT apparatuses, the X-ray tube and the
X-ray detector which are opposed to each other are rotated around
the examinee to collect projection data in many directions, and
these collected projection data are subjected re-constructing
filtering processing for correction of blurring, and then the data
are inversely projected to reconstruct a tomogram(s) of the
examinee.
[0006] The projection data is collected at a discrete projection
angle (hereinafter referred to as "view") of an X-ray beam emitted
from the X-ray tube, and the thus-collected projection data is
referred to as "the projection data at the corresponding view").
The number of views per rotation of the X-ray tube and the X-ray
detector rotating around the examinee, which is required to
reconstruct one CT image, normally ranges from several hundreds to
about one thousand. Furthermore, the projection data of one view
comprises data corresponding to the number of channels.times.the
number of lines of the X-ray detector (in the case of the single
slice type X-ray CT apparatus, the number of lines=1 as described
above).
[0007] Recently, a method of analyzing the composition of an
examinee on the basis of images picked up by irradiating the same
cross-sectional plane with X-ray beams having plural different
energies has been used in the X-ray CT apparatuses as described
above, and this method is called as a multi-energy imaging
method.
[0008] Particularly, when imaging is performed by using two
different kinds of energies, it is called as a dual energy imaging
method.
[0009] A method of measuring tomograms of an average atomic number
and an average density of an examinee by applying the dual energy
imaging method is disclosed in Non-patent Document 1.
[0010] Furthermore, a method of performing the dual energy imaging
while the voltage between the anode and the cathode of the X-ray
tube (hereinafter referred to as "tube voltage") is varied every
X-ray projection angle to change the energy spectrum of the X-ray
(hereinafter referred to as "tube voltage modulation method") is
disclosed in Patent Document 1.
[0011] Non-patent Document 1: R. E. Alvarez and A. Macovski,
"Energy-selective Reconstructions in X-ray Computed Tomography,
"Phys. Med. Biol. Vol. 21, No. 5, pp. 733-744, (1976)
[0012] Patent Document 1: JP-A-10-73544
DISCLOSURE OF THE INVENTION
Problem to be Solved by the Invention
[0013] In order to pick up an image of a fast-moving site of an
examinee such as a heart, coronary artery or the like, the rotating
speed, that is, the scan speed of an imaging system comprising a
pair of an X-ray tube and an X-ray detector has been recently
promoted to increase in an X-ray CT apparatus for medical
application, and X-ray CT apparatuses of about 0.33 to 0.4
[second/revolution] have been practically used.
[0014] As described above, in the X-ray CT apparatus whose scan
speed is increased, the number of views per rotation of the imaging
system in the normal X-ray CT imaging is equal to about 1,000, and
when the scan speed is equal to 0.33 second, imaging is once
carried out per rotation of 0.36.degree. (=360.degree./1,000).
[0015] As described above, when the dual energy imaging is applied
to the conventional X-ray CT apparatus as described above by using
the technique of the patent document 1, the following problems
occur.
[0016] (1) Current flowing between the anode and the cathode of the
X-ray tube (hereinafter referred to as "tube current") is fixed
irrespective of the magnitude of the tube voltage.
[0017] This is because it is difficult to switch the tube current
at high speed due to thermal inertia of the filament temperature of
the X-ray tube and thus the tube current is controlled to be fixed
at a high energy (high tube voltage) and at a low energy (low tube
voltage), so that the tube current is not switched.
[0018] Normally, the X-ray amount absorbed in an examinee increases
when imaging is carried out at a low tube voltage, and thus the
tube current is required to be larger than that at the high tube
voltage.
[0019] Furthermore, when the tube current runs short at the low
tube voltage, the quantum noise in the pickup image increases, and
the quality of the pickup image is lowered.
[0020] Still furthermore, it is possible to obtain sufficient
projection data at the high tube voltage, and thus it is desired
that the tube current is set to be less than that at the low tube
voltage from the viewpoint of reduction of exposure.
[0021] For the foregoing reason, it is originally desired to vary
the tube current (reduce the tube current/increase the tube
current) in connection with the variation of the tube voltage.
[0022] (2) In the technique described in the patent document 1, as
described in paragraph number [0058] of the patent document 1, the
number of views which are obtained in each of the high energy case
and the low energy case during one rotation of a scanner is equal
to 600 and this is a small number. Accordingly, when each view data
obtained in each of the high energy case and the low energy case is
subjected to image reconstruction, there is concern about
occurrence of a radial artifact.
[0023] This is because a delay occurs in a process of supplying the
tube voltage from a power source portion to the X-ray tube due to
parasitic impedance (parasitic inductance and parasitic
electrostatic capacitance) possessed by wires and thus increase of
the switching speed of the tube voltage is restricted, thereby
limiting the number of views.
[0024] The present invention has been implemented in view of the
foregoing problem, and has an object to provide an X-ray generating
apparatus that can switch X-ray beams of high energy and low energy
to each other at high speed in multi-energy imaging using a tube
voltage modulation method, and an X-ray CT apparatus that can
obtain a multi-energy pickup image having high image quality at
high speed by using the X-ray generating apparatus.
Means of Solving the Problem
[0025] In order to attain the above object, there is provided an
X-ray generating apparatus comprising an X-ray tube for generating
and alternately switching X-rays having different energies based on
electron beams generated from plural cathodes by opening and
closing grids, and X-ray control means for controlling the X-rays
having the different energies. The X-rays having the different
energies are emitted while alternatively switched to each other
every adjacent views, thereby obtaining projection data in the
X-ray generating apparatus, and an image is reconstructed from the
projection data concerned. Specifically, the X-ray generating
apparatus is implemented by the following means.
[0026] That is, according to the present invention, in an X-ray
generating apparatus including an X-ray tube for emitting X-ray,
tube current control means for controlling tube current of the
X-ray tube, and X-ray control means for controlling high-energy
X-ray and low-energy X-ray by the tube voltage control means, the
X-ray tube has plural cathodes each of which has a filament, an
anode opposed to the plural cathodes, and grid electrodes each of
which is individually provided every cathode to control discharge
of an electron beam emitted from the cathode, and comprises grid
voltage generating means for generating a voltage to be applied to
each grid electrode, and electron beam emission control means for
alternately applying the grid voltage generated in the grid voltage
generating means to each of the grid electrodes to control the
emission of the electron beam.
[0027] The X-ray tube may be an X-ray tube for emitting plural
electron beams from the plural filaments and forming plural X-ray
focal points on the anode so that the X-ray focal points are spaced
from one another at a predetermined distance on the anode.
Furthermore, the X-ray tube may be an X-ray tube having electron
beam deflecting means which has a deflection coil provided between
the anode and the plural cathodes and deflection current supply
means for supplying the deflection coil with current for deflecting
the directions of the electron beams and deflects the directions of
the electron beams generated from the plural filaments.
[0028] The tube voltage control means has first tube voltage
control means for controlling a high tube voltage corresponding to
the high energy X-ray, and second tube voltage control means for
controlling a low tube voltage corresponding to the low energy
X-ray, and the tube current control means has first tube current
control means for controlling tube current corresponding to the
high energy X-ray, and second tube current control means for
controlling tube current corresponding to the low energy X-ray.
[0029] The tube current controlled by the second tube current
control means is larger than the tube current controlled by the
first tube current control means.
[0030] According to the present invention, in an X-ray CT apparatus
that has a multi-energy imaging function, uses the X-ray generating
apparatus described above and comprises an X-ray tube for
irradiating X-ray with an examinee, an X-ray detector for detecting
X-ray transmitted through the examinee, scanner rotating means
rotating around the examinee while the X-ray tube and the X-ray
detector are mounted therein, X-ray control means for irradiating
X-rays having plural different energies emitted from the X-ray tube
to the same slice position of the examinee while switching the
X-rays every projection angle, and image reconstructing means for
reconstructing projection data detected by the X-ray detector to
obtain a CT image, the X-ray tube comprises plural cathodes each of
which has a filament, an anode opposed to the plural cathodes, and
grid electrodes each of which is individually provided every
cathode to control emission of an electron beam emitted from the
cathode, and the X-ray control means comprises tube current control
means for heating the cathode filaments of the X-ray tube and
control the tube current flowing between the anode and the
cathodes, tube voltage control means for controlling a tube voltage
to be applied between the anode and the cathodes, grid voltage
generating means for generating a voltage to be applied to each
grid electrode, and electron beam emission control means for
applying the grid voltages generated in the grid voltage generating
means to the respective grid electrodes while alternately switching
the grid voltages every projection angle, thereby controlling
emission of the electron beams.
[0031] The X-ray tube of the X-ray CT apparatus may be an X-ray
tube for emitting plural electron beams from the plural filaments
and forming plural X-ray focal points on the anode so that the
X-ray focal points are spaced from each other at a predetermined
distance. Furthermore, the X-ray tube may be an X-ray tube having
electron beam deflecting means that has a deflection coil provided
between the anode and the plural cathodes and deflection current
supply means for supplying the deflection coil with current for
deflecting the directions of the electron beams, and deflects the
directions of the electron beams generated from the plural
filaments.
[0032] The tube voltage control means of the X-ray CT apparatus has
first tube voltage control means for controlling a high tube
voltage corresponding to the high energy X-ray, and second tube
voltage control means for controlling a low tube voltage
corresponding to the low energy X-ray, and the tube current control
means has first tube current control means for controlling tube
current corresponding to the high energy X-ray, and second tube
current control means for controlling tube current corresponding to
the low energy X-ray.
[0033] The tube current controlled by the second tube current
control means is increased to be larger than the tube current
controlled by the first tube current control means, whereby the
image quality is enhanced by reduction of exposure and reduction of
quantum noise.
[0034] The number of projection angles for detecting the projection
data is set to be larger than that under normal CT imaging, thereby
setting the image quality of the normal CT image to the same level
as prior arts.
EFFECT OF THE INVENTION
[0035] According to the present invention, there can be provided
the X-ray generating apparatus that uses the X-ray tube having the
plural cathodes, the anode for forming the X-ray focal points by
electron beams emitted from the filaments of the plural cathodes
and the plural grid electrodes corresponding to the plural cathode
to control the emission of the plural electron beams, and controls
the voltage applied to the grid electrodes so that the high-energy
X-ray and the low-energy X-ray can be switched to each other at
high speed.
[0036] According to the X-ray generating apparatus, the high energy
X-ray and the low energy X-ray are alternately switched and emitted
every adjacent views to obtain projection data. Therefore, the
number of views can be set to the double number of that under
normal CT imaging, and no radial artifact occurs. Therefore, the
enhancement of the image quality can be performed, and issues of a
human body can be clearly discriminated.
[0037] Furthermore, when high-energy X-ray is generated by a high
tube voltage, the tube current is reduced from the viewpoint of
reduction of exposure, and when low energy X-ray is generated by a
low tube voltage, X-ray is generated while the tube current is
increased to the extent that the quantum noise in the image is not
increased. Therefore, both the enhancement of the image quality and
the reduction of the exposure can be attained.
BRIEF DESCRIPTION OF THE DRAWINGS
[0038] FIG. 1 is a schematic diagram showing an X-ray CT apparatus
to which the present invention is applied.
[0039] FIG. 2 is a diagram showing the overall construction of the
X-ray CT apparatus to which the present invention is applied.
[0040] FIG. 3 is a diagram showing the relation between X-ray
irradiation and the construction of an X-ray detector of the X-ray
CT apparatus to which the present invention and X-ray
irradiation.
[0041] FIG. 4 is a diagram showing the construction of an X-ray
tube used in a first embodiment of the present invention.
[0042] FIG. 5 is a diagram showing the structure of the X-ray tube
used in the first embodiment of the present invention.
[0043] FIG. 6 is an enlarged view of the neighborhood of a cathode
of the X-ray embodiment used in the first embodiment of the present
invention.
[0044] FIG. 7 is a diagram showing the detailed construction of the
cathode of the X-ray tube used in the first embodiment of the
present invention.
[0045] FIG. 8 is a diagram showing an example of an electron beam
orbit when an actual focal point is formed on the anode of the
X-ray tube used in the first embodiment of the present
invention.
[0046] FIG. 9 is a diagram showing the positional relationship
between an X-ray focal point position on a scanner rotational plate
for collecting projection data and an X-ray detector in the first
embodiment of the present invention.
[0047] FIG. 10 is a diagram showing the construction of an image
processing apparatus of an X-ray CT apparatus to which the present
invention is applied.
[0048] FIG. 11 is a diagram showing the construction of an X-ray
control device of the X-ray generating apparatus according to the
present invention.
[0049] FIG. 12 is a diagram showing the construction of a tube
voltage generator and a tube voltage controller in the X-ray
control device of the X-ray generating apparatus according to the
present invention.
[0050] FIG. 13 is a diagram showing the construction of a filament
heater and a tube current controller in the X-ray control device of
the X-ray generating apparatus according to the present
invention.
[0051] FIG. 14 is a diagram showing the construction of a grid
voltage generator and a grid switching unit in the X-ray control
device of the X-ray generating apparatus according to the present
invention.
[0052] FIG. 15 is a flowchart showing the operation of a material
identifying function of the X-ray CT apparatus in the first
embodiment of the present invention.
[0053] FIG. 16 is a diagram showing the construction of an X-ray
tube used in a second embodiment of the present invention.
[0054] FIG. 17 is a diagram showing the positional relationship
between an X-ray focal position on a scanner rotational plate for
collecting projection data and an X-ray detector in the second
embodiment of the present invention.
[0055] FIG. 18 is a diagram showing the construction of an X-ray
tube used in a third embodiment of the present invention.
DESCRIPTION OF REFERENCE NUMERALS
[0056] 1 scanner gantry, 4 operation console, 5 operation device, 6
display device, 8 X-ray control device, 9 X-ray tube, 9a tube
voltage generator, 9b filament heater, 9c grid voltage generator,
9d controller, 9d1 tube voltage controller, 9d2 tube current
controller, 9d3 grid opening/closing controller, 9e grid switching
unit, 12 X-ray detecting device, 15 data collecting device, 16
scanner rotating plate, 20 system control device, 21 image
processing device, 21a frame memory, 21b projection data adder, 21c
image reconstructing unit, 21d data converter, 21e material
identifier, 200a first cathode, 200b second cathode, 201a first
grid electrode, 201b second grid electrode, 202a first electron
beam, 202b second electron beam, 203 X-ray focal point, 204
rotational anode, 300 X-ray focal point position on scanner
rotating plate, 600 deflection coil, 701 outer envelope of X-ray
tube, 702 filament, 702a first filament, 702b second filament,
.beta. distance between cathode and anode, .theta. angle of cathode
opposed to anode
BEST MODES FOR CARRYING OUT THE INVENTION
[0057] Preferred embodiments of an X-ray generating apparatus and
an X-ray CT apparatus for performing multi-energy imaging by using
the X-ray generating apparatus according to the invention will be
described in detail with reference to the accompanying
drawings.
First Embodiment
[0058] FIG. 1 is a diagram showing the construction of an X-ray CT
apparatus to which the present invention is applied, and FIG. 2 is
a block diagram showing the overall construction of the X-ray CT
apparatus. The X-ray CT apparatus shown in FIG. 1 irradiates an
examinee with X-ray to collect transmitted X-ray data of the
examinee, and subjects the collected X-ray data to reconstructing
calculation to obtain a tomogram, and it is constructed by a
scanner gantry 1, a table 3 having a top plate 2 on which the
examinee is mounted, and an operation console 4.
[0059] An opening portion 7 in which the examinee is inserted is
provided at the center portion of the scanner gantry 1, and the
table 3 is disposed at the front side of the gantry 1.
[0060] The height of the table 3 can be adjusted by electric
operation, and the top plate 2 is provided on the upper surface of
the table 3. The top plate 2 is constructed to be slidable with
respect to the gantry 1 by electric operation in order to position
the examinee to an imaging position.
[0061] An operation device 5 such as a keyboard, a mouse, etc., and
a display device 6 for displaying various kinds of information such
as patient information, imaging conditions, etc. and a pickup
tomogram are disposed on the operation console 4. An image
processing device described later and a system control device (CPU)
for controlling the overall system are mounted in the operation
console 4. The gantry 1 and the table 2 are controlled by the
system control device (CPU).
[0062] As shown in FIG. 2, the scanner gantry 1 has an X-ray tube 9
which is controlled by the X-ray control device 8 to generate
X-ray, a collimator 11 for narrowing down X-ray emitted from the
X-ray tube 9 to a predetermined irradiation field and an X-ray
detector 12. The X-ray emitted from the X-ray tube 9 is shaped to a
pyramidal X-ray beam, that is, a cone beam by the collimator 11
which is controlled by the collimator control device 10, and
applied to the examinee 23. The X-ray transmitted through the
examinee 23 is incident to the X-ray detector 12.
[0063] The X-ray control device 8 controls the tube voltage and the
tube current so that the tube voltage and the tube current become
the tube voltage and the tube current corresponding to a scan
condition set by the operation device 5. Therefore, a tube
voltage/tube current detecting device 13 for detecting the tube
voltage applied between the anode and cathode of the X-ray tube 9
and the current flowing between the anode and cathode of the X-ray
tube 9 is provided.
[0064] The X-ray detector 12 has plural X-ray detecting elements 14
arranged two-dimensionally in a channel direction and a line
direction as shown in FIG. 3.
[0065] This X-ray detector 14 is constructed by the combination of
a scintillator and a photodiode, for example, and it constitutes an
X-ray incident face which is designed in a cylindrical plane shape
as a whole or curved like a broken line with respect to the channel
direction. The single slice type X-ray detector is not excluded
from this invention.
[0066] Here, a spread angle of the cone beam X-ray in the channel
direction, that is, a fan angle is represented by .alpha., and a
spread angle in the line direction, that is, a cone angle is
represented by .gamma..
[0067] A data collecting device 15 is connected to the
thus-constructed X-ray detector 12, and the data collecting device
15 collects detection data of the X-ray detecting elements 14
constituting the X-ray detector 12.
[0068] The constituent elements from the X-ray control device 8
till the data collecting device 15 are mounted on a rotational
plate 16 (scanner rotating means) of the scanner gantry 1. Driving
force from a rotational plate driving device 18 controlled by a
rotation control device 17 is transmitted by a driving force
transmission system 19, whereby the rotational plate 16 rotates
round the examinee 23.
[0069] The examinee table 3 has an examinee table control device
3a, an examinee table up-and-down moving device 3b and a top plate
moving device 3c. The examinee table control device 3a controls the
examinee table up-and-down moving device 3b so as to set the height
of the table to a proper height, and also controls the top plate
moving device 3c to move the top plate 2 in front-and-rear
direction so that the examinee 23 is fed into and from the X-ray
irradiation space (opening portion) of the scanner gantry 1. A top
plate position sensor 3d detects the top plate position in the body
axial direction and the vertical direction, and controls the top
plate moving device 3c and the examinee table up-and-down moving
device 3b on the basis of the detection information so that the
examinee table control device 3a is set to a correct top plate
position.
[0070] In the thus-constructed scanner gantry 1, when the examinee
23 mounted on the top plate 2 of the examinee table 3 is fed into
the opening portion 7 of the scanner gantry 1 and then the examinee
23 is irradiated with a cone beam X-ray whose cone angle .gamma. is
adjusted by the opening width of the collimator 11, an X-ray image
of the examinee 23 irradiated with the cone beam X-ray is projected
onto the X-ray detector 12, and X-ray transmitted through the
examinee 23 is detected by the X-ray detector 12.
[0071] The operation console 4 has a system control device (CPU) 20
for controlling the overall CT system having a multi-energy imaging
function of the X-ray CT apparatus according to this invention, and
the scanner gantry 1 and the examinee table 3 are connected to the
system control device (CPU) 20. That is, the X-ray control device
8, the collimator control device 10, the data collecting device 15
and the rotation control device 17 in the scanner gantry 1 and the
examinee table control device 3a in the examinee table 3 are
controlled by the system control device (CPU) 20.
[0072] The transmitted X-ray detection data collected by the data
collecting device 15 is taken into the image processing device 21
under the control of the system control device (CPU) 20.
[0073] This image processing device 21 conducts various kinds of
correction processing on the detection data of plural views
collected by the data collecting device 15 to generate projection
data, and performs CT image reconstruction by using this projection
data.
[0074] A scanogram image required for set a scan condition, various
kinds of data, programs for implementing the function of the X-ray
CT apparatus, etc. are stored in a storage device 22 connected to
the system control device (CPU) 20. The CT image reconstructed in
the image processing device 21 is also stored in the storage device
22. Furthermore, the operation device 5 and the display device 6
are connected to the system control device (CPU) 20.
[0075] The operation device 5 is used by an operator to input
various kinds of instructions and information, an image
reconstruction mode, etc. into the system control device (CPU) 20,
and interactively operate the X-ray CT apparatus by using the
display device 6.
[0076] The display device 6 displays the reconstructed image output
from the image processing device 21 and various kinds of
information to be treated by the system control device (CPU)
20.
[0077] The system control device (CPU) 20 predetermines the scan
condition by using an operation instruction input with the operator
operating the operation device 5 and the scanogram image read out
from the storage device 22 before scan is started. That is, the
scanogram image read out from the storage device 22 is displayed on
the display device 6, and the operator indicates the coordinate of
a CT image reconstruction position (hereinafter referred to as
slice position) on the displayed examinee scanogram image by using
the operation device 5, whereby the slice position can be set.
There is a case where the slice position is set by using no
scanogram image. The slice position information set here is stored
in the storage device 22, and also used to set an X-ray condition
(tube voltage, tube current).
[0078] Power to the respective elements mounted in the
thus-constructed gantry 1 is supplied from a commercial alternating
power source by power transmitting means comprising a brush fixed
to a fixed portion of the scanner gantry 1 (not shown) and a
slipping brush provided to a rotating portion (rotational plate 16)
of the scanner gantry 1, and transmission of the transmitted X-pray
detection data collected by the data collecting device 15 to the
image processing device 21 is performed by the same slipping
mechanism as described above or optical signal transmitting
means.
[0079] According to this invention, high energy X-ray and low
energy X-ray from the X-ray tube 9 are alternately switched to each
other every projection angle (view) to perform dual energy imaging
by the thus-constructed X-ray CT apparatus. The X-ray tube shown in
FIG. 4 is used in the first embodiment of the present
invention.
[0080] In FIG. 4, the X-ray tube 9 has a first cathode 200a for
generating an electron beam 202a for generating the high energy
X-ray, a second cathode 200b for generating an electron beam 202b
for generating the low energy X-ray, a first grid 201a
corresponding to the first cathode 200a and a second grid 201b
corresponding to the second cathode 200b which are used to
alternately switch the electron beam 202a and the electrode beam
202b to each other, and a rotational anode 204 on which the
electron beams 202a and 202b emitted from the cathodes 200a and
200b are focused to form the same focal points 203. The focal
points 203 are formed at the same position by adjusting the
distance between the cathode 200a,200b and the rotational anode 204
and the angle of the anode 200a, 200b opposed to the anode 204.
[0081] FIG. 5 shows an example of the structure of thus-constructed
X-ray tube 9, FIG. 6 is an enlarged view of the neighborhood of the
cathode 200a, 200b of the X-ray tube 9, FIG. 7 shows an example of
the detailed structure of the cathode, and FIG. 8 shows an example
of an electron beam orbit when an actual focal point is formed at
the position of the focal point 203.
[0082] As shown in FIG. 5, the X-ray tube 9 comprises the cathodes
200a and 200b for emitting electron beams, the rotational anode 204
opposed to these cathodes and an outer envelope 701 in which the
cathodes 200a and 200b and the rotational anode 204 are
vacuum-tightly sealed. As shown in FIG. 6, the cathodes 200a and
200b are disposed so as to be electrically insulated from each
other, and the focal points 203 are formed at the same position by
adjusting the distance .beta. between the cathode 200a, 200b and
the rotational anode 204 and the angle .theta. of the cathode 200a,
200b opposed to the anode 204.
[0083] Furthermore, as shown in FIG. 7, each of the cathodes 200a
and 200b comprises a filament 702 for emitting an electron beam (a
first filament 702a in the first cathode 200a and a second filament
702b in the second cathode 200b described later), and a focusing
member 703 (a focusing member 703a of the cathode 200a and a
focusing member 703b of the cathode 200b described later) having a
focusing groove portion for focusing the electron beam 202a, 202b
from the filament 702 (see FIG. 4), and the filament 702 is
disposed so as to be electrically insulated from the focusing
member 703 and the grid 201 (the grid 201a corresponding to the
cathode 200a, the grid 201b corresponding to the cathode 200b).
[0084] When the first grid 201a is opened, the electron beam 220a
is emitted from the first filament 702a, and accelerated and
focused by potential gradient between the first cathode 200a and
the rotational node 204 as shown in FIG. 8(a), so that a actual
focal point is formed at the focal point position 203. Likewise,
when the second grid 201b is opened, the electron beam 202b is
emitted from the second filament 702b, and accelerated and focused
by potential gradient between the second cathode 200b and the
rotational anode 204, so that a actual focal point is formed at the
focal point position 203 as shown in FIG. 8(b).
[0085] The opening/closing of the first grid 201a and the second
grid 201b is performed by controlling the voltage applied to the
electrodes of the grids. When the grid is opened, the voltage
between the electrode of the grid and the cathode is set to zero,
and when the grid is closed, a negative voltage of several kV is
applied between the electrode of the grid and the cathode.
[0086] FIG. 9 is a diagram showing the positional relationship
between the X-ray detector 12 and the X-ray focal point position
300 on the rotational plate in connection with the rotation of the
rotational plate 16 when the high energy X-ray and the low energy
X-ray are irradiated while alternately switched with each other
every adjacent projection angles, thereby collecting projection
data.
[0087] In FIG. 9, the position on the rotational plate of the X-ray
focal point 203 at each imaging time every projection angle is
represented by 300(1), 300(2), 300(3), 300(4), . . . , and the
position of the X-ray detector 3 at this time is represented by
12(1), 12(2), . . . . O represents the rotational center of the
rotational plate 16. In this example, the high energy X-ray is
emitted at odd-number positions 300(1), 300(3), 300(5), . . . on
the rotational plate of the X-ray focal point 203, and the low
energy X-ray is emitted at even-number positions 300(2), 300(4),
300(6), . . . . The positions of the X-ray detector 12
corresponding to these X-ray emission positions are represented by
odd numbers and even numbers, and they are represented by a solid
line and a dashed line.
[0088] By the above construction, the high energy X-ray and the low
energy X-ray are prevented from being emitted at the same time.
When the X-ray focal point 203 is located at the position 300(1) of
the rotational plate 16, the examinee 23 is irradiated with the
high energy X-ray, and the X-ray transmitted through the examinee
is detected by the X-ray detector 12 located at the position of
12(1) represented by the solid line of FIG. 9.
[0089] When the rotational plate 16 rotates and thus the X-ray
focal point 203 reaches the position 300(2) of the rotational plate
16, the examinee 23 is irradiated with the low energy X-ray, and
the X-ray transmitted through the examinee is detected by the X-ray
detector 12 located at the position of 12(2) represented by the
dash line of FIG. 9.
[0090] As described above, the high energy X-ray and the low energy
X-ray are alternately irradiated every adjacent projection angles,
the X-rays transmitted through the examinee are detected at the
positions of the X-ray detector 12 corresponding to the projection
angles concerned, the detected data are collected by the data
collecting device 15, the collected data are transmitted to the
image processing device 21, and a dual energy pickup image is
generated in the image processing device 21.
[0091] FIG. 10 is a block diagram showing a part of generating the
dual energy pickup image by using the data collected in the data
collecting device 15.
[0092] In FIG. 10, the image processing device 21 comprises a data
correcting unit 21a for correcting (offset correction, gain
correction, etc.) the X-ray detection data collected in the data
collecting device 15, a memory 21b for storing projection data
corrected in the data correcting unit 21a, a data reading unit 21c
for reading out projection data based on the high-energy X-ray
imaging and projection data based on the low energy X-ray imaging
stored in the memory 21b according to a reconstructing mode
instruction input to the operation device 5 by the operator, a data
adder 21e for adding pairs of adjacent views of the projection data
of the high energy X-ray imaging and the projection data of the low
energy X-ray imaging (300(1) and 300(2), 300(3) and 300(4), . . .
in FIG. 3) read out from the memory 21b, an image reconstructing
unit 21f for conducting well-known reconstructing calculation such
as filtering, back projection, etc. by using the projection data
read out from the memory 21b according to the reconstructing mode
instruction to thereby reconstruct a CT image of the examinee 23, a
data converter 21d for reading out the projection data based on the
high energy X-ray imaging and the projection data based on the low
energy X-ray imaging in the adjacent views which are stored in the
memory 21b, and creating an X-ray attenuation image based on the
photoelectric effect (hereinafter referred to as photoelectric
effect image) and an X-ray attenuation image based on the Compton
scattering (hereinafter referred to as Compton image) by using a
well-known method disclosed in the non-patent Document 1, and a
material identifier 21g for determining an average atomic number
and an average density of the examinee by using the converted data
obtained in the data converter 21d on the basis of the CT images of
the photoelectric effect image and the Compton image according to
the well-known method disclosed in the Non-patent document 1,
identifying material for each pixel in the CT image on the basis of
the information of the average atomic number and the average
density, and outputting the information to the display device
6.
[0093] In FIG. 10, the reconstructing mode input to the operation
device 5 can be selected from
(1) a separate mode in which the projection data based on the high
energy X-ray imaging and the projection data based on the low
energy X-ray imaging are individually subjected to image
reconstruction, (2) an image addition mode in which the data adder
21e is operated, and data obtained by adding (or subtracting) the
projection data based on the high energy X-ray imaging and the
projection data based on the low energy X-ray imaging as described
above is subjected to image reconstruction, and (3) a material
identifying mode for identifying material as described
elsewhere.
[0094] A dedicated processor, a well-known general-purpose
processor or the like is applied to the data adder 21e, the image
reconstructing unit 21f, the data converter 21d and the material
identifier 21g.
[0095] Furthermore, it is desired that the number of projection
angles at which the high energy X-ray and the low energy X-ray are
irradiated, that is, the number of views is larger than that of the
normal CT imaging to suppress the radial artifact and discriminate
the issues of a human body, and typically it is preferably twice as
large as 1024 views under the normal CT imaging, that is, 2048
views.
[0096] As described above, in the case of the view number of 2048,
when one-rotation time of the rotational plate, that is, one-scan
time is set to 0.33 second, the high energy X-ray and the low
energy X-ray are alternately applied to the examinee 23 to perform
imaging every time the rotational plate 16 rotates at 0.17 [deg],
and thus it is required to perform the switching operation between
the high energy X-ray and the low energy X-ray at high speed.
[0097] FIG. 11 is an example of a block diagram showing the
construction of the X-ray control device 8 (FIG. 2) for performing
the dual energy imaging while alternately switching the high energy
X-ray and the low energy X-ray from the X-ray tube 9.
[0098] The X-ray control device 8 comprises a tube voltage
generator 9a for generating a high tube voltage applied between the
anode 204 and the first cathode 200a and a low tube voltage applied
between the anode 204 and the second cathode 200b, a filament
heater 9b for heating the first filament 702a and the second
filament 702b so as to obtain the tube current corresponding to the
high tube voltage and the low tube voltage, a grid voltage
generator 9c (grid voltage generating means) for generating a
voltage for opening/closing the first grid 201a and the second grid
201b, a grid switching unit 9e (electron beam emission control
means) for switching the first grid 201a and the second grid 201b,
and a controller 9d for controlling the tube voltage generator 9a,
the filament heater 9b, the grid voltage generator 9c and the grid
switching unit 9e.
[0099] The controller 9d comprises a tube voltage controller 9d1
for controlling the tube voltage generator 9a to output the high
tube voltage and the low tube voltage, a tube current controller
9d2 for controlling the filament current so that the filament
current is equal to the tube current set in accordance with the
high tube voltage and the low tube voltage, and a grid voltage
controller 9d3 for controlling the grid voltage generating unit 9c
so that the output voltage of the grid voltage generating unit 9c
is equal to zero (open the grid) and a high voltage of about -1000V
(close the grid).
[0100] When the high tube voltage set value and the low tube
voltage set value are set and input from the operation device 5 of
the operation console 4 to the system control device (CPU) 20, the
tube voltage set signals corresponding to these set values are
generated in a tube voltage/tube current set signal generator 20a,
and the tube voltage controller 9d1 performs the control so that
the generated high tube voltage set signal and low tube voltage set
signal are coincident with the actual high tube voltage and low
tube voltage of the X-ray tube 9 which are detected by the tube
voltage/tube current detecting device 13 (FIG. 2).
[0101] FIG. 12 shows an example of the circuit construction of the
tube voltage generator 9a and the tube voltage controller 9d1 (tube
voltage control means).
[0102] In FIG. 12, the tube voltage controller 9d1 comprises a
converter circuit 9d11 for converting an AC voltage of a
three-phase AC power source 25 to a DC voltage, a smoothing
capacitor for smoothing the output DC voltage (three-phase
full-wave rectified voltage) of the converter circuit 9d11, a first
inverter circuit 9d13 for converting the DC voltage of the
smoothing capacitor 9d12 to an AC voltage having a frequency (for
example, 20 kHz or more) which is further higher than the
three-phase AC power source frequency and also controlling the high
tube voltage, a second inverter circuit 9d14 for converting the DC
voltage of the smoothing capacitor 9d12 to an AC voltage having a
high frequency (for example, 20 kHz or more) and controlling the
low tube voltage, and a tube voltage control signal generator 9d15
for generating the high tube voltage control signal of the first
inverter circuit 9d13 and the low tube voltage control signal of
the second inverter circuit 9d14. The tube voltage control signal
generator 9d15 generates a high tube voltage control signal and a
low tube voltage control signal for controlling the operating
phases of the first inverter circuit 9d13 and the second inverter
circuit 9d14 so that the tube voltage set signal (the high tube
voltage set signal and the low tube voltage set signal) generated
in the tube voltage/tube current set signal generator 20a (see FIG.
11) of the system control device (CPU) 20 is coincident with the
tube voltage detection value detected by the tube voltage/tube
current detecting device 13.
[0103] The tube voltage generator 9a comprises a high tube voltage
generator for generating a high tube voltage and a low tube voltage
generator for generating a low tube voltage. The high tube voltage
generator has is equipped with a first high voltage transformer 9a1
having two secondary windings for boosting the output voltage of
the first inverter circuit 9d13 to the high tube voltage, and high
voltage rectifying circuits 9a2 and 9a3 for converting the voltages
of the two secondary windings of the first high voltage transformer
9a1 to DC voltages. The high voltage rectifying circuits 9a2 and
9a3 are connected to each other in series, and the connection point
thereof is grounded.
[0104] On the other hand, the low tube voltage generator has a
second high voltage transformer 9a4 for boosting the output voltage
of the second inverter circuit 9d14, and a high voltage rectifying
circuit 9a5 for rectifying the secondary winding voltage of the
second high voltage transformer 9a4 to a DC voltage, and the low
tube voltage is generated from the output voltage of the high
voltage rectifying circuit 9a2 and the output voltage of the high
voltage rectifying circuit 9a5. In this case, the positive DC
output terminal of the high voltage rectifying circuit 9a5 is
grounded.
[0105] The positive DC output terminal of the high voltage
rectifying circuit 9a2 is connected to the rotational anode 204 of
the X-ray tube 9, the negative DC output terminal of the high
voltage rectifying circuit 9a3 is connected to the first cathode
200a of the X-ray tube 9, and the negative DC output terminal of
the high voltage rectifying circuit 9a5 is connected to the second
cathode 200b of the X-ray tube 9.
[0106] For example when the first inverter circuit 9d13 and the
second inverter circuit 9d14 are controlled by the thus-constructed
tube voltage control means (the first tube voltage control means,
the second tube voltage control means) so that the output voltage
of the high voltage rectifying circuit 9a2 is equal to 70 kV, the
output voltage of the high voltage rectifying circuit 9a3 is equal
to -70 kV and the output voltage of the high voltage rectifying
circuit 9a5 is equal to -10 kV, the high tube voltage of 140 kV
{=70 kV-(-70 kV)} is applied between the anode 204 and the first
cathode 200a of the X-ray tube 9 and the low tube voltage of 80 kV
{=70 kV-(-10 kV)} is applied between the anode 204 and the second
cathode 200b of the X-ray tube 9 because the outer envelope 701 of
the X-ray tube 9 is grounded.
[0107] In this case, the first filament 702a and the second
filament 702b are controlled to be heated by the filament heater 9b
so as to obtain the tube current corresponding to the high tube
voltage and the low tube voltage.
[0108] FIG. 13 shows an example of the circuit construction of the
filament heater 9b and the tube current controller 9d2 (tube
current control means).
[0109] In FIG. 13, the filament heater 9b comprises a first
filament heater 9b1 for heating the first filament 702a, a second
filament heater 9b2 for heating the second filament 702b and a DC
power source 9b3.
[0110] The first filament heater 9b1 comprises a first filament
heating inverter circuit 9b11 for converting the voltage of the DC
power source 9b3 to an AC voltage having a high frequency and also
variably controlling the filament current for heating the first
filament 702a of the X-ray tube 9, and a first filament heating
transformer 9b12 for insulting the output of the first filament
heating inverter circuit 9b11 and supplying the output to the first
filament 702a. Likewise, the second filament heating circuit 9b2
comprises a second filament heating inverter circuit 9b21 for
converting the voltage of the DC power source 9b3 to an AC voltage
having a high frequency and also variably controlling the filament
current for heating the second filament 702b of the X-ray tube 9,
and a second filament heating transformer 9b22 for insulating the
output of the second filament heating inverter circuit 9b21 and
supplying the output to the second filament 702b.
[0111] The tube current controller 9d2 generates a first filament
current control signal for controlling the filament current of the
first filament 702a corresponding to the low tube current and a
second filament current control signal for controlling the filament
current of the second filament 702b corresponding to the high tube
current, which are used to control the operation phases of the
first filament heating inverter circuit 9b11 and the second
filament heating inverter circuit 9b21 so that the tube current set
signal (the low tube current corresponding to the high tube voltage
and the high tube current set signal corresponding to the low tube
voltage) generated in the tube voltage/tube current set signal
generator 20a of the system control device (CPU) 20 is coincident
with the tube current detection value detected by the tube
voltage/tube current detecting device 13.
[0112] The first filament 702a and the second filament 702b are
heated by the filament heater 9b and the tube current controller
9d2 in advance so that preset tube current is obtained (first tube
current control means, second tube current control means), the set
tube current can be made to flow simultaneously with the
opening/closing of the first grids 201a and 201b, so that the
problem caused by the thermal inertia of the filament temperature
in the prior arts can be solved, and the tube current can be
switched at high speed.
[0113] FIG. 14 shows an example of the circuit construction of the
grid voltage generator 9c and the grid switching unit 9e.
[0114] In FIG. 14, the grid voltage generator 9c comprises a first
grid voltage generator 9c1 for generating a negative high voltage
of about 1000V for closing the first grid 201a of the X-ray tube 9,
and a second grid voltage generator 9c2 for generating a negative
high voltage of about 1000V for closing the second grid 201b (grid
voltage generating means), and the grid switching unit 9e (electron
beam emission control means) comprises a first grid opening/closing
switch 9e1 for controlling the opening/closing of the first grid
201a, and a second grid opening/closing switch 9e2 for controlling
the opening/closing of the second grid 201b.
[0115] The grid switching unit 9e is supplied with a grid switching
opening/closing signal generated in the grid switching signal
generator 20b (see FIG. 11) of the system control device (CPU) 20
through the grid opening/closing controller 9d3 of the controller
9d. When the first grid 201a is opened, the grid switching unit 9e
connects a of the first grid opening/closing switch 9e1 to c so
that the voltage applied to the first grid 201a is set to zero, and
when the first grid 201a is closed, the grid switching unit 9e
connects a of the first grid opening/closing switch 9e1 to b so
that a negative voltage of about 1000V is applied to the first grid
201a.
[0116] Likewise, on the basis of the grid switching opening/closing
signal, the grid switching unit 9e connects d of the second grid
opening/closing switch 9e2 to f so that the voltage applied to the
second grid 201b is set to zero when the second grid 201b is
opened, and connects d of the second grid opening/closing switch
9e2 to e so that a negative voltage of about 1000V is applied to
the second grid 201b when the second grid 201b is closed.
[0117] In order to switch the opening/closing switch at high speed,
the first grid opening/closing switch 9e1 and the second grid
opening/closing switch 9e2 is implemented by using a semiconductor
switch disclosed in "High-speed Pulse See-through System (Takano,
others)", Japan Radiation Technique Society Magazine vol. 57, No.
10, FIG. 2 Structure of MOSEFT super-cascade high-voltage
semiconductor switching module (October in 2001), or switching
control means based on an optical signal disclosed in FIGS. 8 to 11
of JP-A-2003-317996, for example.
[0118] By the thus-constructed grid voltage generator 9c and grid
switching unit 9e, the first grid 201a and the second grid 201b are
alternately switched to each other on the basis of the grid
switching signal generated in the grid switching signal generator
20b, whereby the high energy X-ray and the low energy X-ray can be
generated on the basis of the electron beam generated from the
cathodes corresponding to the first and second grids.
[0119] Next, the operation of the thus-constructed X-ray CT
apparatus will be described with reference to the flowchart of FIG.
15.
[0120] (1) Start (S10)
[0121] The operator turns on the power source switch of the
operation console 4 to start imaging (multi-energy CT imaging).
[0122] (2) Set Scan Conditions (S11)
[0123] First, the operator sets scan conditions by using the
operating device 5 of the operation console 4 before the
imaging.
[0124] The main scan conditions contain an X-ray condition A and an
X-ray condition B for generating two kinds of X-ray energies to be
emitted from the X-ray tube 9, a scan speed (the rotational speed
of the rotational plate 16), a slice position of the examinee 23, a
slice range, etc, and also a reconstructing mode, etc. as described
above.
[0125] The X-ray condition contains an X-ray condition A (high
energy X-ray) based on a high tube voltage and tube current with
which the radiation dosage of the examinee corresponding the high
tube voltage is as small as possible, and an X-ray condition B (low
energy X-ray) based on a low tube voltage and tube current which is
larger than the tube current of the X-ray condition A to the extent
that the quantum noise in the pickup image does not increase and
thus the image quality of the pickup image is not degraded.
[0126] (3) Start Scan (Imaging) (S12)
[0127] Subsequently, the operator starts rotation of the rotational
plate 16 by operating the operating device 5 of the operation
console 4. Driving force from the rotational plate driving device
18 controlled by the rotation control device 17 is transmitted to
the rotational plate 16 by the driving force transmission system
19, whereby the rotational plate 16 rotates around the examinee 23.
Simultaneously with the start of the rotation of the rotational
plate 16, the tube voltages corresponding to the X-ray conditions A
and B are respectively applied between the rotational anode 204 and
the first cathode 200a of the X-ray tube 1 and between the
rotational anode 204 and the second cathode 200b of the X-ray tube
1, and also the filaments 702a and 702b are heated so that the
temperatures of the filaments 702a and 702b are equal to the
filament temperatures corresponding to the tube current of the
X-ray conditions A and B. In order to prevent electron beams from
being emitted from these filaments, the first grid 201a and the
second grid 201b are electrically closed.
[0128] Then, at the time point when the rotational speed of the
rotational plate 16 reaches the set scan speed (for example, the
rotational speed corresponding to one rotation/0.33 second), the
first grid 201a is opened to irradiate the examinee 23 with the
X-ray of the X-ray condition A, thereby starting the imaging.
[0129] (4) Collect Projection Data (S13)
[0130] When the imaging is started, the system control device (CPU)
20 alternately opens/closes the first grid 201a and the second grid
201b every projection angle of the rotational plate 16
corresponding to each of the first grid 201a and the second grid
201b (every 0.17[degree] when the number of views is set to 2048
and the one-scan time is set to 0.33 second, for example)
Accordingly, the X-ray based on the X-ray condition A and the X-ray
based on the X-ray condition B are alternately applied to the
examinee 23.
[0131] The X-ray transmitted through the examinee 23 in this
imaging operation is detected by the X-ray detector 12,
successively collected as projection data in the data collecting
device 15, and then transmitted to the image processing device
21.
[0132] The data transmitted to the image processing device 21 are
subjected to various kinds of correction processing in the data
correcting unit 21a, and then these projection data are
successively recorded in the memory 21b.
[0133] The system control device (CPU) 20 performs the processing
of S21 to S24, the processing of S31 to S34 or the processing of
S41 to S45 according to the reconstructing mode which is input and
set to the operation console 4 by the operator in S11, the image
reconstruction and the display of the processing result on the
display device 6.
[0134] (5) When the Separate Mode is Selected
[0135] When the separate mode is selected in S11, the system
control device (CPU) 20 executes each of the following steps S21 to
S24.
(5-1) Separate and Read Out Projection Data (S21)
[0136] When this mode is selected, the system control device (CPU)
20 instructs the data reading unit 21c to read out the projection
data of the high energy X-ray imaging views 300(1), 300(3), 300(5),
. . . and the projection data of the low energy X-ray imaging views
300(2), 300(4), 300(6), . . . as individual pairs from the memory
21b. Accordingly, the projection data (scan data) based on the high
energy X-ray imaging and the projection data (scan data) based on
the low energy X-ray imaging data are successively transmitted to
the image reconstructing unit 21f.
(5-2) Reconstruct Image (S22, S23)
[0137] The image reconstructing unit 21f successively executes the
image reconstruction on the projection data of the high energy
X-ray imaging successively supplied from the memory 21b and the
image reconstruction on the projection data of the low energy X-ray
imaging, and creates two CT images (a first CT image and a second
CT image). These CT images are stored into the storage device 22 in
the console device 4.
(5-3) Display Image (S24)
[0138] The two reconstructed CT images are displayed on the screen
of the display device 6, and supplied for doctor's image diagnosis.
A system of selectively displaying two images or arranging two
images in parallel on the screen may be applied as an image display
style. Such a technique is well known in this technical field, and
thus the description thereof is omitted.
[0139] According to the separate mode described above, an internal
organ or tissue in the examinee can be imaged according to the
X-ray absorption characteristic. Furthermore, the two CT images
thus obtained are in the same time phase, and thus the diagnosis
based on the multi-energy CT image of a moving internal organ can
be performed in a short time and with less exposure.
[0140] (6) When the Image Addition Mode is Selected
[0141] When the image addition mode is selected in S11, the system
controller (CPU) 20 executes each of the following steps S31 to
S34.
(6-1) Read Out Projection Data (S31)
[0142] When this mode is selected, the system controller (CPU) 20
instructs the data reading unit 21c to successively read out the
data from the memory 21b while the high energy X-ray imaging view
and the low energy X-ray imaging view which are adjacent to each
other are paired like a pair of 300(1) and 300(2), a pair of 300(3)
and 300(4), a pair of 300(5) and 300(6), . . . . Accordingly, the
pairs of the projection data are successively transmitted to the
data adder 21e.
(6-2) Create Composite Projection Data (S32)
[0143] The data adder 21e executes addition processing (any one of
simple addition, addition average and subtraction) on the paired
data of the transmitted projection data to create composite
projection data. The number of views of the composite projection
data is set to the half of the total views of the number of the
high energy X-ray imaging views and the number of the low energy
X-ray imaging views, and these data are transmitted to the image
reconstructing unit 21f.
(6-3) Reconstruct Image (S33)
[0144] The image reconstructing unit 21f executes the
reconstructing calculation on the basis of the composite projection
data transmitted from the data adder 21e to crate an CT image
(third CT image). The thus-created third CT image is stored into
the storage device 22 in the console device 4.
(6-4) Display Image (S34)
[0145] The reconstructed third CT image is displayed on the screen
of the display device 6, and supplied for doctor's image
diagnosis.
[0146] According to the image addition mode described above, there
can be obtained an image having compensated contrast resolution
which is short when the examinee is imaged by only one of the high
energy X-ray imaging and the low energy X-ray imaging.
[0147] The image addition mode can be also implemented by
subjecting the two CT image obtained in the separate mode to
addition processing, addition averaging processing or the
subtraction processing.
[0148] (7) When the Material Identifying Mode is Selected
[0149] When the material identifying mode is selected in S11, the
system controller (CPU) 20 executes each of the following steps S41
to S45.
(7-1) Generate Data for Reconstructing Photoelectric Effect Image
and Compton Image (S41)
[0150] When this mode is selected, the system controller (CPU) 20
reads out the projection data picked up by the high energy X-ray
and the projection data picked up by the low energy X-ray at two
adjacent views which are recorded in the memory 21b, and creates,
in the data converter 21d, data for reconstructing a photoelectric
effect image based on the photoelectric effect and a Compton image
based on Compton scattering by using the well-known method
disclosed in the non-patent document 1.
(7-2) Reconstruct Photoelectric Effect Image and Compton Image
(S42)
[0151] Respective CT images of the photoelectric effect image and
the Compton image are reconstructed in the image reconstructing
unit 21f by using the reconstructing data of the photoelectric
effect image and the Compton image which are generated in the data
converter 21d (fourth CT image and fifth CT image).
(7-3) Reconstruct CT Image Having Average Atomic Number and Average
Density (S43)
[0152] ON the basis of the reconstructed CT images of the
photoelectric effect image and the Compton image, a CT image (sixth
CT image) having an average atomic number and an average density of
the examinee 23 is reconstructed in the reconstructing device 21f
by using the well-known method disclosed in the non-patent document
1.
(7-4) Identify Material (S44)
[0153] The material identification is performed on the basis of the
reconstructed sixth CT image of the average atomic number and the
average density in the material identifier 21g (material
identifying means) by applying a well-known method to each
pixel.
(7-5) Display Identification Result (S45)
[0154] The identification result obtained in S44 is displayed on
the screen of the display device 6. As a display style of the
identification result, window information provided with hue
information which is different every tissue or material is
displayed, and also the hues corresponding to the window
information are provided to sites which are different in tissue,
material in the sixth CT image.
[0155] In this case, by changing the color every material, plural
identification results can be displayed at the same time. Typical
examples of the materials to be identified contain bone tissue,
lung tissue, muscle and fat of a human body, contrast material,
etc.
[0156] (8) Finish (S50)
[0157] When the processing of each mode is finished, the operator
turns off the power source switch of the console device 4. At the
same time when the imaging (scan) of the indicated imaging range is
finished, the system control device (CPU) 20 instructs the X-ray
control device 8 and the rotational plate control device to finish
the imaging, thereby finishing he radiation of X-ray from the X-ray
tube 1 and stopping the rotation of the rotational plate 16.
[0158] Furthermore, the above imaging operation and the respective
kinds of processing (the image reconstruction processing, the
addition processing, the data creation processing of the
photoelectric effect image and the Compton image, the material
identification processing) are executed in parallel to one another,
and the results are successively displayed on the display device
6.
[0159] As described above, according to the first embodiment, there
is provided the X-ray generating apparatus that controls the
voltages to be applied to the grid electrodes and switches the high
energy X-ray and the low energy X-ray to each other at high speed
by using the X-ray tube having the two cathodes, the anode for
forming one X-ray focal point by electron beams emitted from the
filaments of the two cathodes and the two grid electrodes
corresponding to the two cathodes for controlling emission of the
plural electron beams. By the X-ray generating apparatus, the high
energy X-ray and the low energy X-ray are irradiated while
alternately switched to each other every adjacent views, thereby
obtaining the projection data. Therefore, the number of views can
be set to the double of the number of views under the normal CT
imaging, and also no radial artifact occurs, so that the image
quality of the image can be enhanced, and the tissues of the human
body can be clearly discriminated from one another.
[0160] In this case, the number of views are set to the double of
that under the normal CT imaging, and thus the image quality of the
CT image can be set to the same level as the prior arts.
[0161] Furthermore, when the high energy X-ray based on the high
tube voltage is generated, the X-ray is generated while the tube
current is reduced from the viewpoint of reducing the exposure, and
when the low energy X-ray based on the low tube voltage is
generated, the X-ray is generated while the tube current is
increased to the extent that the quantum noise in the image does
not increase. Therefore, the enhancement of the image quality and
the reduction of the exposure can be attained.
[0162] Furthermore, the first tube voltage control means and the
second tube voltage control means can independently control the
high tube voltage and the low tube voltage respectively, and thus
the dual energy imaging based on the combination of any high energy
X-ray and low energy X-ray can be performed.
Second Embodiment
[0163] According to a second embodiment of the present invention,
the dual energy imaging is performed by using an X-ray tube 9'
shown in FIG. 16.
[0164] In FIG. 16, the X-ray tube 9' has a first anode 200a' for
generating an electron beam 202a' for generating high energy X-ray,
a second cathode 200b' for generating an electron beam 202b' for
generating low energy X-ray, a first grid 201a' corresponding to
the first cathode 200a' and a second grid 201b' corresponding to
the second cathode 200b' for alternately switching the electron
beam 202a' and the electron beam 202b' to each other, and a
rotational anode 204' for forming two focal points of a first X-ray
focal point 203a' and a second X-ray focal point 203b' by focusing
the electron beams 202a' and 202b' emitted from the cathodes 200a'
and 200b'. The first focal point 203a' and the second focal point
203b' are formed so as to be spaced from each other at a distance
of d by adjusting the distance between the cathode 200a', 200b' and
the rotational anode 204' and the angles of the cathodes 200a' and
200b' with respect to the rotational anode 204'.
[0165] 702a' represents a first filament for generating the
electron beam 202a', and 702b' represents a second filament for
generating the electron beam 202b'.
[0166] The thus-constructed X-ray tube 9' is different in only the
distance between the cathode 200a', 200b' and the rotational anode
204' and the angles of the cathodes 200a' and 200b' with respect to
the rotational anode 204' to form the first and second two focal
points 203a' and 203b', and the structure of the X-ray tube 9', the
enlarged view of the neighborhood of the cathodes 200a' and 200b'
of the X-ray tube 9', the detailed structure of the cathodes and
the electron orbit when the actual focal points are formed at the
positions of the focal points 203a' and 203b' are substantially the
same as FIGS. 5, 6, 7 and 8 of the first embodiment, and thus the
description thereof is omitted.
[0167] FIG. 17 is a diagram showing the positional relationship
between the X-ray detector 12 and the X-ray focal point on the
rotational orbit in connection with the rotation of the rotational
plate 16 when the high energy X-ray and the low energy X-ray are
irradiated from the two focal points 203a' and 203b' of the X-ray
tube 9' by using the X-ray tube 9' of FIG. 16 while alternately
switched to each other every adjacent projection angles, thereby
collecting projection data.
[0168] In FIG. 17, the position on the rotational orbit of the
first X-ray focal point 203a' in each imaging operation at each
projection angle is represented by 500(1), 500(3), 500(5), . . . ,
and the position of the X-ray detector 12 at this time is
represented by 12(1), 12(3), . . . .
[0169] The position on the rotational orbit of the second X-ray
focal point 203b' is represented by 500(2), 500(4), 500(6), . . . ,
and the position of the X-ray detector 12 is represented by 12(2),
12(4), . . . . In FIG. 16, only 12(1) and 12(2) are shown, and
these are represented by a solid line and a dashed line.
[0170] In the second embodiment, the high energy X-ray is emitted
from the first X-ray focal point 203a' at the odd-number positions
500(1), 500(3), 500(5), . . . on the rotational orbit, the
rotational plate 16 rotates and the low energy X-ray is emitted
from the X-ray focal point 203b' spaced from the first X-ray focal
point 203a' at a distance d at even-number positions 500(2),
500(4), 500(6), . . . . Corresponding to 500(1)+d, 500(3)+d,
500(5)+d, . . . on the rotational plate 16. The transmitted X-ray
of the examinee is detected at the positions 12(1), 12(2), . . . .
Of the X-ray detector 12 which correspond to the above X-ray
emission positions.
[0171] As in the case of the first embodiment, the high energy
X-ray and the low energy X-ray are emitted every adjacent
projection angles, the X-ray transmitted through the examinee is
detected at the position of the X-ray detector 12 corresponding to
the projection angle concerned, and the detected data are collected
in the data collecting device 15 and transmitted to the image
processing device 21 to generate a dual energy pickup image in the
image processing device.
[0172] By the above construction, the same effect as the first
embodiment can be obtained, and the X-ray can be surely emitted
from the second focal point 203b' at the position spaced at the
distance d. Therefore, thereby preventing reduction of the
resolution which is concerned to occur due to the positional
displacement between the position of 500(2), 500(4), 500(6), . . .
and the X-ray focal point position.
Third Embodiment
[0173] A third embodiment of the present invention performs the
dual energy imaging by using an X-ray pipe 9'' shown in FIG.
18.
[0174] The difference from the X-ray pipe 9' of the second
embodiment resides in that electron beam deflecting means
comprising a deflection coil 600 for deflecting the direction of an
electron beam emitted from the cathode and a magnetic field
generating current supply source (deflection current supply means)
(not shown) for making current flow through the coil to generate
magnetic field.
[0175] In FIG. 18, the X-ray tube 9' has a first cathode 200a' for
generating an electron beam 601a for generating high energy X-ray,
a second cathode 200b' for generating an electron beam 601b for
generating low energy X-ray, a first grid 201a' corresponding to
the first cathode 200a' and a second grid 201b' corresponding to
the second cathode 200b' for alternately switching the electron
beam 601a and the electron beam 601b to each other, a deflection
coil 600 for changing the travel directions of the electron beams
601a, 601b, and a rotational anode 204' for focusing the electron
beams 601a and 601b emitted from the cathodes 200a' and 200b' and
forming two focal points 602a and 602b of the first X-ray focal
point 602a and the second X-ray focal point 602b.
[0176] 702a' represents a first filament for generating the
electron beam 202a', and 702b' represents a second filament for
generating the electron beam 202b'.
[0177] The deflection coil 600 can generate magnetic field in the
vertical direction of the sheet plane, and control the intensity of
the magnetic field to adjust the travel directions of the electron
beams 601a, 601b in the up-and-down direction of the sheet plane.
Particularly by applying the magnetic field to the electron beam
601a and the electron beam 601b in the opposite directions
respectively, the travel directions of the respective electron
beams can be changed to the opposite directions to each other. That
is, by adjusting the intensity of the magnetic field and the
applying direction of the magnetic field, the positions of the
first X-ray focal point 602a and the second X-ray focal point 602b
which are formed by the electron beams 601a and 601b, and the
distance between these focal points can be arbitrarily
adjusted.
[0178] Of course, it is possible to set the distance d to zero so
that the focal point position is single.
[0179] It is more preferable that the distance d between the focal
points is changed in accordance with the scan speed. Accordingly,
in such a case, for example, the distance d may be variably
controlled by the electron beam deflecting means so that the
distance d is increased when the scan speed is high.
[0180] As in the case of the first embodiment, the high energy
X-ray and the low energy X-ray are emitted every projection angle,
the X-ray transmitted through the examinee is detected at the
positions of the X-ray detector 12 corresponding to the projection
angles, the detected data are collected by the data collecting
device 15, and the collected data are transmitted to the image
processing device 21 to generate a dual energy pickup image in the
image processing device.
[0181] In the dual energy X-ray imaging using the thus-constructed
X-ray tube 9'', the same effect as the first embodiment and the
second embodiment can be obtained. In addition, as shown in FIG.
17, the distance d of the focal point position which is required to
be shifted on an arc 501 is varied in accordance with the
projection number per rotation, however, there is an advantage that
the focal point position can be varied in accordance with various
projection numbers by using the X-ray tube 9'' shown in FIG.
18.
[0182] The first to third embodiments in which the dual energy
imaging is performed by alternately switching the high energy X-ray
and the low energy X-ray to each other every adjacent views by
using the X-ray generating apparatus for switching the high energy
X-ray and the low energy X-ray has been described. However, the
present invention is not limited to the dual energy imaging, and
the present invention may be applied to multi-energy based on three
or more X-ray energies.
[0183] The following X-ray generating apparatuses of (1), (2) and
(3) are applied under the multi-energy imaging.
[0184] (1) An X-ray generating apparatus comprising an X-ray tube
having an anode for forming one focal point by electron beams from
three or more plural cathodes and plural grids corresponding to the
cathodes for switching the electron beams from the plural cathodes,
and plural X-ray energy control means for controlling plural X-ray
energies generated from the X-ray tube.
[0185] (2) An X-ray generating apparatus comprising an X-ray tube
having an anode for forming three or more X-ray focal points by
electron beams from three or more plural cathodes, and plural grids
corresponding to the cathodes for switching the electron beams from
the plural cathodes, and plural X-ray energy control means for
controlling plural X-ray energies generated from the X-ray
tube.
[0186] (3) An X-ray generating apparatus an X-ray tube which is
obtained by providing the X-ray tube of (2) with electron beam
deflecting means for deflecting the direction of the electron beam
and deflects the direction of the electron beam by the electro beam
deflecting means to form plural X-ray focal points, and plural
X-ray energy control means for controlling plural X-ray energies
generated from the X-ray tube.
[0187] In the above embodiments, the high energy X-ray and the low
energy X-ray are switched every adjacent views to perform the dual
energy imaging by one scan. However, subsequently to the above
scan, the imaging based on X-ray energy different from that of the
previous scan may be performed on the same slice position.
[0188] The X-ray generating apparatus according to this invention
and the X-ray CT apparatus using the X-ray generating apparatus are
applied to a human body as a target. However, the present invention
is not limited to the human body, but applicable to a baggage
scanner for detecting existence of detonating powder, that is,
explosive material in baggage.
* * * * *